Cone-beam computerized tomography with a flat-panel imager

ABSTRACT

A radiation therapy system that includes a radiation source that moves about a path and directs a beam of radiation towards an object and a cone-beam computer tomography system. The cone-beam computer tomography system includes an x-ray source that emits an x-ray beam in a cone-beam form towards an object to be imaged and an amorphous silicon flat-panel imager receiving x-rays after they pass through the object, the imager providing an image of the object. A computer is connected to the radiation source and the cone beam computerized tomography system, wherein the computer receives the image of the object and based on the image sends a signal to the radiation source that controls the path of the radiation source.

[0001] Applicants claim, under 35 U.S.C. §119(e), the benefit ofpriority of the filing date of Feb. 18, 2000, of U.S. Provisional PatentApplication Serial No. 60/183,590, filed on the aforementioned date, theentire contents of which are incorporated herein by reference.

BACKGROUND OF THE INVENTION

[0002] 1. Field of the Invention

[0003] The present invention relates generally to a cone-beam computedtomography system and, more particularly, to a cone-beam computedtomography system that employs an amorphous silicon flat-panel imagerfor use in radiotherapy applications where the images of the patient areacquired with the patient in the treatment position on the treatmenttable.

[0004] 2. Discussion of the Related Art

[0005] Radiotherapy involves delivering a prescribed tumorcidalradiation dose to a specific geometrically defined target or targetvolume. Typically, this treatment is delivered to a patient in one ormore therapy sessions (termed fractions). It is not uncommon for atreatment schedule to involve twenty to forty fractions, with fivefractions delivered per week. While radiotherapy has proven successfulin managing various types and stages of cancer, the potential exists forincreased tumor control through increased dose. Unfortunately, deliveryof increased dose is limited by the presence of adjacent normalstructures and the precision of beam delivery. In some sites, thediseased target is directly adjacent to radiosensitive normalstructures. For example, in the treatment of prostate cancer, theprostate and rectum are directly adjacent. In this situation, theprostate is the targeted volume and the maximum deliverable dose islimited by the wall of the rectum.

[0006] In order to reduce the dosage encountered by radiosensitivenormal structures, the location of the target volume relative to theradiation therapy source must be known precisely in each treatmentsession in order to accurately deliver a tumorcidal dose whileminimizing complications in normal tissues. Traditionally, a radiationtherapy treatment plan is formed based on the location and orientationof the lesion and surrounding structures in an initial computerizedtomography or magnetic resonance image. However, the location andorientation of the lesion may vary during the course of treatment fromthat used to form the radiation therapy treatment plan. For example, ineach treatment session, systematic and/or random variations in patientsetup (termed interfraction setup errors) and in the location of thelesion relative to surrounding anatomy (termed interfraction organmotion errors) can each change the location and orientation of thelesion at the time of treatment compared to that assumed in theradiation therapy treatment plan. Furthermore, the location andorientation of the lesion can vary during a single treatment session(resulting in intrafraction errors) due to normal biological processes,such as breathing, peristalsis, etc. In the case of radiation treatmentof a patient's prostate, it is necessary to irradiate a volume that isenlarged by a margin to guarantee that the prostate always receives aprescribed dose due to uncertainties in patient positioning and dailymovement of the prostate within the patient. Significant dose escalationmay be possible if these uncertainties could be reduced from currentlevels (˜10 mm) to 2-3 mm.

[0007] Applying large margins necessarily increases the volume of normaltissue that is irradiated, thereby limiting the maximum dose that can bedelivered to the lesion without resulting in complication in normalstructures. There is strong reason to believe that increasing the dosedelivered to the lesion can result in more efficacious treatment.However, it is often the case that the maximum dose that can be safelydelivered to the target volume is limited by the associated dose tosurrounding normal structures incurred through the use of margins.Therefore, if one's knowledge of the location and orientation of thelesion at the time of treatment can be increased, then margins can bereduced, and the dose to the target volume can be increased withoutincreasing the risk of complication in normal tissues.

[0008] A number of techniques have been developed to reduce uncertaintyassociated with systematic and/or random variations in lesion locationresulting from interfraction and intrafraction errors. These includepatient immobilization techniques (e.g., masks, body casts, bite blocks,etc.), off-line review processes (e.g., weekly port films,population-based or individual-based statistical approaches, repeatcomputerized tomography scans, etc.), and on-line correction strategies(e.g., pre-ports, MV or kV radiographic or fluoroscopic monitoring,video monitoring, etc.).

[0009] It is believed that the optimum methodology for reducinguncertainties associated with systematic and/or random variations inlesion location can only be achieved through using an on-line correctionstrategy that involves employing both on-line imaging and guidancesystem capable of detecting the target volume, such as the prostate, andsurrounding structures with high spatial accuracy.

[0010] An on-line imaging system providing suitable guidance has severalrequirements if it is to be applied in radiotherapy of this type. Theserequirements include contrast sensitivity sufficient to discernsoft-tissue; high spatial resolution and low geometric distortion forprecise localization of soft-tissue boundaries; operation within theenvironment of a radiation treatment machine; large field-of-view (FOV)capable of imaging patients up to 40 cm in diameter; rapid imageacquisition (within a few minutes); negligible harm to the patient fromthe imaging procedure (e.g., dose much less than the treatment dose);and compatibility with integration into an external beam radiotherapytreatment machine.

[0011] Several examples of known on-line imaging systems are describedbelow. For example, strategies employing x-ray projections of thepatient (e.g., film, electronic portal imaging devices, kVradiography/fluoroscopy, etc.) typically show only the location of bonyanatomy and not soft-tissue structures. Hence, the location of asoft-tissue target volume must be inferred from the location of bonylandmarks. This obvious shortcoming can be alleviated by implantingradio-opaque markers on the lesion; however, this technique is invasiveand is not applicable to all treatment sites. Tomographic imagingmodalities (e.g., computerized tomography, magnetic resonance, andultrasound), on the other hand, can provide information regarding thelocation of soft-tissue target volumes. Acquiring computerizedtomography images at the time of treatment is possible, for example, byincorporating a computerized tomography scanner into the radiationtherapy environment (e.g., with the treatment table translated betweenthe computerized tomography scanner gantry and the radiation therapygantry along rails) or by modifying the treatment machine to allowcomputerized tomography scanning. The former approach is a fairlyexpensive solution, requiring the installation of a dedicatedcomputerized tomography scanner in the treatment room. The latterapproach is possible, for example, by modifying a computer tomographyscanner gantry to include mechanisms for radiation treatment delivery,as in systems for tomotherapy. Finally, soft-tissue visualization of thetarget volume can in some instances be accomplished by means of anultrasound imaging system attached in a well-defined geometry to theradiation therapy machine. Although this approach is not applicable toall treatment sites, it is fairly cost-effective and has been used toillustrate the benefit of on-line therapy guidance.

[0012] As illustrated in FIGS. 1(a)-(c), a typical radiation therapysystem 100 incorporates a 4-25 MV medical linear accelerator 102, acollimator 104 for collimating and shaping the radiation field 106 thatis directed onto a patient 108 who is supported on a treatment table 110in a given treatment position. Treatment involves irradiation of alesion 112 located within a target volume with a radiation beam 114directed at the lesion from one or more angles about the patient 108. Animaging device 116 may be employed to image the radiation field 118transmitted through the patient 108 during treatment. The imaging device116 for imaging the radiation field 118 can be used to verify patientsetup prior to treatment and/or to record images of the actual radiationfields delivered during treatment. Typically, such images suffer frompoor contrast resolution and provide, at most, visualization of bonylandmarks relative to the field edges.

[0013] Another example of a known on-line imaging system used forreducing uncertainties associated with systematic and/or randomvariations in lesion location is an X-ray cone-beam computerizedtomography system. Mechanical operation of a cone beam computerizedtomography system is similar to that of a conventional computerizedtomography system, with the exception that an entire volumetric image isacquired through a single rotation of the source and detector. This ismade possible by the use of a two-dimensional (2-D) detector, as opposedto the 1-D detectors used in conventional computerized tomography. Thereare constraints associated with image reconstruction under a cone-beamgeometry. However, these constraints can typically be addressed throughinnovative source and detector trajectories that are well known to oneof ordinary skill in the art.

[0014] As mentioned above, a cone beam computerized tomography systemreconstructs three-dimensional (3-D) images from a plurality oftwo-dimensional (2-D) projection images acquired at various angles aboutthe subject. The method by which the 3-D image is reconstructed from the2-D projections is distinct from the method employed in conventionalcomputerized tomography systems. In conventional computerized tomographysystems, one or more 2-D slices are reconstructed from one-dimensional(1-D ) projections of the patient, and these slices may be “stacked” toform a 3-D image of the patient. In cone beam computerized tomography, afully 3-D image is reconstructed from a plurality of 2-D projections.Cone beam computerized tomography offers a number of advantageouscharacteristics, including: formation of a 3-D image of the patient froma single rotation about the patient (whereas conventional computerizedtomography typically requires a rotation for each slice); spatialresolution that is largely isotropic (whereas in conventionalcomputerized tomography the spatial resolution in the longitudinaldirection is typically limited by slice thickness); and considerableflexibility in the imaging geometry. Such technology has been employedin applications such as micro-computerized tomography, for example,using a kV x-ray tube and an x-ray image intensifier tube to acquire 2-Dprojections as the object to be imaged is rotated, e.g., through 180° or360°. Furthermore, cone beam computerized tomography has been usedsuccessfully in medical applications such as computerized tomographyangiography, using a kV x-ray tube and an x-ray image intensifier tubemounted on a rotating C-arm.

[0015] The development of a kV cone-beam computerized tomography imagingsystem for on-line tomographic guidance has been reported. The systemconsists of a kV x-ray tube and a radiographic detector mounted on thegantry of a medical linear accelerator. The imaging detector is based ona low-noise charge-coupled device (CCD) optically coupled to a phosphorscreen. The poor optical coupling efficiency (−10⁻⁴) between thephosphor and the CCD significantly reduces the detective quantumefficiency (DQE) of the system. While this system is capable ofproducing cone beam computerized tomography images of sufficient qualityto visualize soft tissues relevant to radiotherapy of the prostate, thelow DQE requires imaging doses that are a factor of 3-4 times largerthan would be required for a system with an efficient coupling (e.g.−50% or better) between the screen and detector.

[0016] Another example of a known auxiliary cone beam computerizedtomography imaging system is shown in FIG. 2. The auxiliary cone beamcomputerized tomography imaging system 200 replaces the CCD-based imagerof FIGS. 1(a)-(c) with a flat-panel imager. In particular, the imagingsystem 200 consists of a kilovoltage x-ray tube 202 and a flat panelimager 204 having an array of amorphous silicon detectors that areincorporated into the geometry of a radiation therapy delivery system206 that includes an MV x-ray source 208. A second flat panel imager 210may optionally be used in the radiation therapy delivery system 206.Such an imaging system 200 could provide projection radiographs and/orcontinuous fluoroscopy of the lesion 212 within the target volume as thepatient 214 lies on the treatment table 216 in the treatment position.If the geometry of the imaging system 200 relative to the system 206 isknown, then the resulting kV projection images could be used to modifypatient setup and improve somewhat the precision of radiation treatment.However, such a system 200 still would not likely provide adequatevisualization of soft-tissue structures and hence be limited in thedegree to which it could reduce errors resulting from organ motion.

[0017] Accordingly, it is an object of the present invention to generateKV projection images in a cone beam computerized tomography system thatprovide adequate visualization of soft-tissue structures so as to reduceerrors in radiation treatment resulting from organ motion.

BRIEF SUMMARY OF THE INVENTION

[0018] One aspect of the present invention regards a radiation therapysystem that includes a radiation source that moves about a path anddirects a beam of radiation towards an object and a cone-beam computertomography system. The cone-beam computer tomography system includes anx-ray source that emits an x-ray beam in a cone-beam form towards anobject to be imaged and an amorphous silicon flat-panel imager receivingx-rays after they pass through the object, the imager providing an imageof the object. A computer is connected to the radiation source and thecone beam computerized tomography system, wherein the computer receivesthe image of the object and based on the image sends a signal to theradiation source that controls the path of the radiation source.

[0019] A second aspect of the present invention regards a method oftreating an object with radiation that includes moving a radiationsource about a path, directing a beam of radiation from the radiationsource towards an object and emitting an x-ray beam in a cone beam formtowards the object. The method further includes detecting x-rays thatpass through the object due to the emitting an x-ray beam with anamorphous silicon flat-panel imager, generating an image of the objectfrom the detected x-rays and controlling the path of the radiationsource based on the image.

[0020] Each aspect of the present invention provides the advantage ofgenerating KV projection images in a cone beam computerized tomographysystem that provide adequate visualization of soft-tissue structures soas to reduce errors in radiation treatment resulting from organ motion.

[0021] Each aspect of the present invention provides an apparatus andmethod for improving the precision of radiation therapy by incorporatinga cone beam computerized tomography imaging system in the treatmentroom, the 3-D images from which are used to modify current andsubsequent treatment plans.

[0022] Each aspect of the present invention represents a significantshift in the practice of radiation therapy. Not only does thehigh-precision, image-guided system for radiation therapy address theimmediate need to improve the probability of cure through doseescalation, but it also provides opportunity for broad innovation inclinical practice.

[0023] Each aspect of the present invention may permit alternativefractionation schemes, permitting shorter courses of therapy andallowing improved integration in adjuvant therapy models.

[0024] Each aspect of the present invention provides valuable imaginginformation for directing radiation therapy also provides an explicit3-D record of intervention against which the success or failure oftreatment can be evaluated, offering new insight into the means by whichdisease is managed.

[0025] Additional objects, advantages and features of the presentinvention will become apparent from the following description and theappended claims when taken in conjunction with the accompanyingdrawings.

BRIEF DESCRIPTION OF THE DRAWINGS

[0026] FIGS. 1(a)-(c) schematically show the geometry and operation of aconventional radiation therapy apparatus;

[0027]FIG. 2 schematically shows a perspective view of a known radiationtherapy apparatus including an auxiliary apparatus for cone beamcomputerized tomography imaging;

[0028]FIG. 3 is a diagrammatic view of a bench-top cone beamcomputerized tomography system employing a flat-panel imager, accordingto a first embodiment of the present invention;

[0029]FIG. 4 is a schematic illustration of the geometry and proceduresof the cone beam computerized tomography system shown in FIG. 3;

[0030] FIGS. 5(a)-5(d) are graphs depicting the fundamental performancecharacteristics of the flat-panel imager used in the cone beamcomputerized tomography system of FIG. 3;

[0031] FIGS. 6(a)-6(d) show various objects used in tests to investigatethe performance of the cone beam computerized tomography system of thepresent invention, including a uniform water cylinder, six low-contrastinserts in a water bath, a steel wire under tension with a water bath,and an euthanized rat, respectively;

[0032] FIGS. 7(a)-7(d) depict uniformity of response of the cone beamcomputerized tomography system of the present invention, including axialand sagittal slices through volume images of a uniform water bath,radial profiles, and a vertical signal profile, respectively;

[0033] FIGS. 8(a)-8(d) illustrate the noise characteristics of the conebeam computerized tomography system of the present invention, includingaxial and sagittal noise images from volume reconstructions of a uniformwater bath, radial noise profiles, and vertical nose profiles,respectively;

[0034] FIGS. 9(a)-9(b) depict response linearity and voxel noise,respectively, for the cone beam computerized tomography system of thepresent invention and a conventional computerized tomography scanner;

[0035] FIGS. 10(a)-10(c) depict the noise-power spectrum from the conebeam computerized tomography system of the present invention, includinga gray scale plot of the axial noise-power spectrum, the noise-powerspectrum measured at various exposures, and the noise-power spectrum forthe cone beam computerized tomography system compared to a conventionalcomputerized tomography scanner, respectively;

[0036] FIGS. 11(a)-11(b) depict the spatial resolution of the cone beamcomputerized tomography system of the present invention, including thesurface plot of an axial slice image of the thin steel wire shown inFIG. 6(c) and the modulation transfer function measured for the conebeam computerized tomography system and for a conventional computerizedtomography scanner, respectively;

[0037] FIGS. 12(a)-12(b) show images of a low-contrast phantom obtainedfrom the cone beam computerized tomography system of the presentinvention and a conventional computerized tomography scanner,respectively;

[0038] FIGS. 13(a)-13(i) show cone beam computerized tomography imagesof the euthanized rat shown in FIG. 6(d), including regions of the lungs(FIGS. 13(a)-13(c)), the kidneys (FIGS. 13(d)-13(f)), and the lowerspine (FIGS. 13(g) -13(i));

[0039] FIGS. 14(a)-14(d) show volume renderings of cone beamcomputerized tomography images of the euthanized rat shown in FIG. 6(d)illustrating the degree of spatial resolution achieved in delineatingstructures of the vertebra, including volume renderings with axial andsagittal cut planes showing the skeletal anatomy along with soft-tissuestructures of the abdomen, volume renderings with axial and sagittal cutplanes, window to show skeletal features only, a magnified view of aregion of the spine and ribs of the rat, and a magnified view of a partof two vertebra, respectively;

[0040] FIGS. 15(a)-15(b) depict the axial images of euthanized rat shownin FIG. 6(d) obtained from the cone beam computerized tomography systemof the present invention and a conventional computerized tomographyscanner, respectively;

[0041]FIG. 16 is a graph showing detected quantum efficiency calculatedas a function of exposure for an existing and hypothetical flat-panelimager configuration;

[0042] FIGS. 17(a)-(e) are diagrammatic views of several angularorientations of a wall-mounted cone beam computerized tomography systememploying a flat-panel imager, according to a second embodiment of thepresent invention;

[0043]FIG. 18 shows a side view of the cone beam computerized tomographysystem of FIG. 17 when employing a first embodiment of a support for aflat-panel imager according to the present invention;

[0044]FIG. 19(a) shows a perspective exploded view of a mounting to beused with the support for a flat-panel imager of FIG. 18;

[0045]FIG. 19(b) shows a perspective exploded view of a rotationalcoupling to be used with the mounting of FIG. 19(a);

[0046] FIGS. 20(a)-(b) schematically shows a front view of thewall-mounted cone beam computerized tomography system of FIG. 17 whenemploying a second embodiment of a support for a flat-panel imageraccording to the present invention;

[0047] FIGS. 21(a)-(b) schematically shows a front view of thewall-mounted cone beam computerized tomography system of FIG. 17 whenemploying a third embodiment of a support for a flat-panel imageraccording to the present invention;

[0048]FIG. 22 is a diagrammatic view of a portable cone beamcomputerized tomography system employing a flat-panel imager accordingto fifth embodiment of the present invention;

[0049] FIGS. 23(a)-(d) are diagrammatic sketches illustrating thegeometry and operation of the cone beam computerized tomography imagingsystems of FIGS. 17-22;

[0050]FIG. 24 is a flow-chart showing an embodiment of the processesinvolved in acquiring a cone beam computerized image for the cone beamcomputerized tomography imaging systems of FIGS. 17-22;

[0051]FIG. 25 is a perspective drawing illustrating an embodiment of amethod for geometric calibration of the imaging and treatment deliverysystems of FIGS. 17-22; and

[0052]FIG. 26 is a flow-chart showing an embodiment of the processesinvolved in the image-guided radiation therapy systems of FIGS. 17-22,based on cone beam computerized tomography imaging of a patient, on-linecorrection of setup errors and organ motion, and off-line modificationof subsequent treatment plans.

PREFERRED EMBODIMENTS OF THE INVENTION

[0053] A bench-top cone beam computerized tomography (CBCT) system 300is shown in FIG. 3, according to an embodiment of the present invention.The CBCT system 300 was constructed to mimic the geometry of the CBCTscanner currently installed on a linear accelerator, with asource-to-axis distance of 1000 mm and a source-detector distance of1600 mm. The primary components of the system 300 include an x-ray tube302, a rotation stage 304 and flat-panel imager (FPI) 306. Thesecomponents are rigidly mounted to an optical bench 308. The relativeposition of these components is controlled by three translation stages,including an object stage 310, a yobject stage 312 and a yimage stage314, which are used during initial setup to accurately determine andcontrol the imaging geometry. The cone beam computerized tomographysystem 300 generates images of an object 316, identified throughout as aphantom, mounted on the rotation stage 304. Each stage 310, 312 and 314contains a home or limit switch, and the imaging geometry is referencedto the location of these switches with a reproducibility of ±0.01 mm.The specific geometries used in the discussion herein are shown in FIG.4, and are set to simulate the imaging geometry that would beimplemented for a cone beam computerized tomography system incorporatedon a radiotherapy treatment machine. Table 1 below shows the parametersof the system 300.

[0054] A set of alignment lasers 318 allow visualization of the axis ofrotation 320 and the source plane perpendicular to the axis of rotation320 and intersects focal spot 322 of the x-ray source or tube 302. Theaxis of rotation 320 is positioned such that it intersects the centralray 324 between the focal spot 322 and the detector plane 326 (+0.01mm). The flat plane imager 326 is positioned such that the piercingpoint (i.e., the intersection of the central ray and the image plane) iscentered on the imaging array (i.e., between columns #256 and #257,±0.01 mm), with a quarter-pixel offset applied to give improved viewsampling for cone beam computerized tomography acquisitions in which theobject 316 is rotated through 360°. The stage 310 is controlled manuallyby means of a positioning micrometer. The source-to-object (SOD) andsource-to-image (SID) distances were measured to within ±0.5 mm and givean objection magnification of 1.60, equal to that of the imaging systemon the linear accelerator. The cone angle for this geometry is −7.1.

[0055] Radiographic exposures used in the acquisition procedure areproduced under computer control with a 300 kHU x-ray tube 302, such asGeneral Electric Maxi-ray 75 and a 100 kW generator, such as the GeneralElectric MSI-800. The tube 302 has a total minimum filtration of 2.5 mmA1, with an additional filtration of 0.127 mm Cu to further harden thebeam, and a nominal focal spot size of 0.6 mm. The 100 kV beam ischaracterized by first and second HVLs of 5.9 and 13.4 mm A1,respectively. The accelerating potential of the generator was monitoredover a one-week period and was found to be stable to within ±1%. Allexposures were measured using an x-ray multimeter, such as the RTIElectronics, Model PMX-III with silicon diode detector.

[0056] The exposures for the cone beam computerized tomographyacquisitions are reported in terms of exposure to air at the axis ofrotation 320 in the absence of the object 316. The same method ofreporting exposure can be used for the images acquired on theconventional scanner. For the conventional scanner, the exposure perunit charge is measured with the gantry rotation disabled and thecollimators set for a 10 mm slice thickness, thereby guaranteeingcomplete coverage of the silicon diode. The exposure per unit charge at100 kVp was 9.9 mR/mAs and 14.9 mR/mAs for the bench-top andconventional scanners, respectively.

[0057] The flat panel imager 306 can be the EG&G Heimann Optoelectronics(RID 512-400 AO) that incorporates a 512×512 array of a-Si:H photodiodesand thin-film transistors. The electromechanical characteristics of theimager are shown in Table 1. The flat plane imager 306 is read-out atone of eight present frame rates (up to 5 frames per second) andoperates asynchronously of the host computer 328 schematically shown inFIG. 4. The analog signal from each pixel is integrated by ASICamplifiers featuring correlated double-sampling noise reductioncircuitry. Digitization is performed at 16 bit resolution. The valuesare transferred via an RS-422 bus to a hardware buffer in the hostcomputer 328. The processor in the host computer 328 is interrupted whena complete frame is ready for transfer to host memory. TABLE 1 CBCTCharacteristic Value Acquisition Geometry Source-axis-distance (S_(AD))103.3 cm Source-imager-distance (S_(ID)) 165.0 cm Cone angle 7.1°Maximum angular rotation rate 0.5° /sec Field of view (FOV) 12.8 cmX-ray Beam/Exposure Characteristics Beam energy 100 kVp Added filtration1.5 mm A1 + 0.129 mm Cu Beam quality HVL₁ = 5.9 MM A1  HVL₂ = 13.4 MM A1Scatter-to-primary ratio 0.18, 1:5 (11 cm object) Frame time 6.4 secTube output at (SAD) 9.34 mR/mAs Exposure rate (at SID) 3.65 mR/mAsFlat-Panel Imager Designation RID 512-400 AO Array format 512 × 512pixels Pixel pitch 400 μm Area ˜20.5 20.5 cm² Pixel fill factor 0.80Photodiode charge capacity ˜62 Pc ASIC amplifier charge capacity ˜23 pCASIC amplifier noise ˜12,700 e ADC bit-depth 16 bit TFT thermal noise(on) ˜1800 e Photodiode Shot Noise (1 fps) ˜1200 e Digitazation noise˜630 e Nominal frame rate 0.16 fps Maximum frame rate 5 fps X-rayconverter 133 mg/cm₂Gd₂O₂S:Tb Acquisition Procedure Number ofprojections 300 Angular increment 1.2° Total rotation angle 360° Maximumangular rotation rate 05 ^(D)/s Reconstruction Parameters Reconstructionmatrix 561 × 561 × (1-512), 281 × 281 × (1-500)  Voxel size  0.25 × 0.25× 0.25 mm2, 0.5 × 0.5 × 0.25   W. parameter 1.60 γ, cutoff frequencymodification 1.0 α, modified Hamming filter parameter 0.50 Range ofconvolution ± 25 mm

[0058] The cone-beam scanning procedure includes a repeated sequence ofradiographic exposure, array readout, and object rotation. The timing ofthis procedure is driven by the asynchronous frame clock of the flatplane imager readout electronics. A conservative frame time of 6.4 s wasused. Between the periodic frame transfers from the flat plane imager306, the host computer advances the motorized rotation stage 304 andtriggers the x-ray generator or tube 302. The rotor of the x-ray tube302 remains spinning throughout the scanning procedure. The controlsoftware allows the operator to specify the number of frames betweenexposures. This was designed as a mechanism to investigate methods ofreducing the amount of lag in sequential projections. The detectorsignal from a group of nine pixels in the bare-beam region of the flatplane imager 306 is monitored to measure and verify the stability ofeach radiographic exposure. Exposures outside tolerance are trapped andrepeated at the same projection angle. Each projection image is writtento hard disk between frame transfer and motor rotation. After theprojections are acquired, a set of flood and dark field images (20 each)are collected to construct gain and offset images for flat-fieldprocessing of the projection images.

[0059] In addition to gain and offset corrections, median filtration(3×3) is performed using a pre-constructed map of unstable pixels.Finally, the signal in each projection is normalized to account forsmall variations in x-ray exposure, this is performed using a cluster ofnine pixels in the periphery of the detector well outside the objectsshadow.

[0060] A volumetric computerized tomography data set is reconstructedfrom the projections using a filtered back-projection technique. Thefilter used in the reconstruction is constructed using Webb'sthree-parameter formula. The parameters and their corresponding valuesare shown in Table 1. In the current configuration, the reconstructionfield of vision is limited to a 12.4 cm diameter cylinder, approximately12.1 cm in length; the lateral extent of objects to be reconstructedmust lie well-within this cylinder. The voxel values in the resultingvolumetric data sets are scaled linearly to produce a mean CT number ofzero in air and 1000 in water. The time required to filter (100 elementkernel) and back-project a single projection (512×512) on to a281×281×500 voxel data set was 1 minute and 21 seconds.

[0061] The basic signal and noise characteristics of the flat planeimager 306 were measured. The detector gain and linearity are presentedin FIG. 5(a). For an x-ray beam energy of 120 kVp, the detector gain wasmeasured to be 18.2×10⁵ e/mR/pixel (17.8×10⁶ e/mR at 100 kVp). Thedetector exhibits excellent linearity with exposure up to 50% of itssensitive range (5 mR). The various additive electronic noise sourcesand their magnitudes are listed in Table 1. The total additiveelectronic noise is found to depend upon frame time, ranging from 13,300e at a frame time of 200 ms to 22,500 e at a frame time of 25.6 s. Theamplifier noise (12,700 e) is the dominant component at high framerates. The significance of amplifier noise on the zero-frequencydetective quantum efficiency (DQE) was studied using a cascaded systemmodel that analyzes signal and noise propagation in the FPI 306.

[0062]FIG. 5(b) shows the dependence of detective quantum efficiency onexposure for the RID 512-400AO, as well as for two hypothetical imagerswith reduced amplifier noise. The primary quantum efficiency for thedetector is approximately 0.57; losses due to energy absorption noiseand additive sources reduce the detective quantum efficiency to ˜0:41for exposures above 1 mR. For exposures below 0.1 mR, the detectivequantum efficiency falls rapidly for amplifier noise values comparableto that found in the EG&G detector. Thus for thicker/denser objects[e.g., a pelvis (˜30 cm water)] resulting in significantly reduced doseto the detector (e.g., ˜0.001 mR) improvements in amplifier noise(and/or x-ray converter, e.g. Csl;TI) will significantly improvedetective quantum efficiency.

[0063] The temporal stability of the detector dark signal is presentedin FIG. 5(c). This plot corresponds to a selected group of ‘typical’pixels. The dark signal drifts significantly during the first 2 h ofoperation, which correlates with the change in temperature within theflat panel imager enclosure. After the temperature has stabilized, thedark signal also stabilizes. Based on these results, all cone beamcomputerized tomography scans were performed after the array had beenpowered-on for at least 2 hours. In some regions of the array, the darksignal does not stabilize, even after thermal equilibrium. It is assumedthat these regions are the result of variations in the arraymanufacturing process.

[0064] The continuously changing scene in computerized tomographynecessitates a detector with rapid read out and minimal temporalblurring, or ‘lag.’ Such characteristics have been measured using ashort, intra-frame, x-ray exposure. FIG. 5(d) shows the pixel signalfollowing a single radiographic exposure applied within the acquisitionperiod of frame number 0. Subsequent frames exhibit lag signal rangingfrom ˜4% to ˜0.04% for frame members 1 through 9. It is interesting andimportant to note that the lag demonstrates a dependence not upon frametime, but also exclusively upon the number of frames.

[0065] Prior to reconstruction, the projections are corrected forstationary pixel-to-pixel variations in offset and gain. Defectivepixels with significant variations in dark field signal or with aberrantsignal response are median filtered. The resulting projections arepadded by an additional 128 columns prior to reconstruction. The valueof the padded pixels is set row-by-row to the average of the 7 pixels atthe periphery of the array. Finally, to account for small variations inx-ray tube output, the signal in each projection is normalized usingsignal measured from the bare-beam monitors pixels mentioned above (ninepixels). The pre-construction processing can be performed on a 250 MHzUltraSparc processor, such as the Enterprise 450, Sun Microsystems,Sunnyvale, Calif.

[0066] Feldkamp's filtered back-projection algorithm can be used toreconstruct the data set. Images are reconstructed on a Cartesian matrixof voxels 561×561×N, where the number of slices, N, depends on theobject of interest. The voxel size used in these reconstructions wastypically 0.25×0.25×0.25 mm. The filtering used in the reconstructionfollows the formalism of Webb. Table 1 contains the three parametersthat specify the filter used in these investigations. Upon completion ofthe reconstruction, an offset and scale parameters are constant for a 9mm set of reconstruction and acquisition parameters. The reconstructionof the volumetric cone beam computerized tomography data sets is alsoperformed on the UltraSparc system.

[0067] The uniformity of response of the imaging system 300 over thethree-dimensional (3-D) field-of-view (FOV) was studied by imaging acylindrical water bath [110 mm diameter]. Scans of the same phantom werealso acquired on the conventional scanner. The response was examinedalong both radial and vertical profiles through the reconstructedvolume.

[0068] The noise in reconstructed images of the water bath was studiedas a function of x-ray exposure. Images were acquired at exposures of131, 261, 653, 1310, 3260, and 6530 mR. The images were reconstructed ona 561×561×11 matrix with voxel dimensions of 0.25 mm on a side. For allreconstructions, the reconstruction filter was fixed at the parametersspecified in Table 1. Varying these parameters can have a significanteffect on the noise characteristics of the reconstructed images. Thenoise characteristics of these image sets were analyzed by analysis ofthe standard deviation in CT number in 5×5×1 regions throughout the dataset, and by calculation of the noise power spectrum (NPS) from the 3Ddata sets. Both methods of analysis were performed as a function ofexposure. The relative stability of the noise was assessed by examiningthe uniformity of the noise over the entire 3-D data set. These resultsindicated that the noise characteristics of the data set vary onlyslightly with location. These initial results lend support to theapplication of noise power analysis, since stability is a necessarycondition for proper interpretation of noise power results.

[0069] The noise-power spectrum (NPS) was analyzed from the volumetricdata by extension of methods employed for analysis of known 2-Dprojection images. The volume data was normalized such that the mean CTnumber within the water cylinder was 1000. A tetragonal region(256×256×20 voxels) within the water cylinder was cropped from thevolume, and a small number of voxel defects (always<1%) were 3×3 medianfiltered. In order to obtain a convergent 2-D central slice of the 3-DFourier transform, the twenty slices were averaged along thez-direction, and it was found that averaging more slices did not affectthe noise-power spectrum, i.e, the data was convergent. A backgroundslice formed from the average of 81 slices in a separate scan wassubtracted in order to reduce background trends. Low-frequency trendswere further reduced by subtraction of a planar fit to the data,yielding a 2-D zero-mean realization. The two-dimensional Fast FourierTransform (FFT) was computed from ensembles of sixteen 64×64non-overlapping regions within the realization, and the results wereaveraged. The results were normalized to account for voxel size and foraverage in z, and the volume under the noise-power spectrum was comparedto the square of the standard deviation. The resulting noise-powerspectrum represents a central slice in the (u_(x)u_(y)) domain, i.e.,the Fourier counterpart to the (x, y) domain. Strips along the u, axiswere extracted in order to show 1-D power spectra, NPS(u_(x)), e.g., arevarious exposure levels.

[0070] The noise characteristics of the cone beam computerizedtomography system 300 were compared to those of the conventionalcomputerized tomography scanner. To allow meaningful comparison, the twosystems must demonstrate identical response over the range of signalvariation. The response was tested by scanning an electron densityphantom (shown in FIG. 6(b)) with the two systems. Seven inserts withcoefficients near that of water were inserted into a 110 mm diameterwater bath. The inserts are taken from the RMI electron density phantomhaving nominal CT numbers. In FIG. 6(b), clockwise from the top: CTSolid Water (CT#1001), BR-SRI Breast (CT#945), BRN-SR2 Brain (CT#1005),C133 Resin Mix (CT#1002), LV1 Liver (CT#1082), and, Polyethylene(CT#897). This phantom was imaged at equivalent exposure and kVp withboth the cone beam computerized tomography system 300 and theconventional scanner.

[0071] The attenuation coefficients (relative to water) reported by thecone beam computerized tomography system 300 were compared to thosereported by the conventional scanner. A first-order fit to the measureddata was calculated to determine the relative linearity of the twosystems. The noise characteristics of the conventional scanner were alsomeasured using the water cylinder test phantom described above imageswere acquired at 100 kVp with a slice thickness of 1 mm at fourdifferent exposure levels (743, 1490, 2970, and 5940 mR). Three imageswere acquired at each exposure level. Reconstructions were performed onthe conventional scanner using the ‘High Res Head (#1 H)’, ‘StandardHead (#2)’, and ‘Smooth Abdomen (#3)’ filters. The noise analysis wasidentical to that applied to the cone beam computerized tomography datasets. In order to compare noise results measured on each system,analysis of the cone beam computerized tomography data sets was repeatedwherein the cone beam computerized tomography data was first averageover 2×2×4 voxels to yield an equivalent (0.5×0.5×1 mm) voxel size tothat given by the conventional scanner.

[0072] The spatial frequency transfer characteristics of the cone beamcomputerized tomography system 300 were measured using a wire testobject, shown in FIG. 6(c). The test object consists of a 0.254 mmdiameter steel wire suspended in a 50 mm diameter water bath. Thephantom was imaged on the cone beam computerized tomography system 300(at 100 kVp) with the wire centered on the axis of rotation 320 and withthe wire located −30 mm off-axis. The resulting images werereconstructed on a high resolution reconstruction grid of 0.1×0.1×0.25mm³ using the filter described in Table 1. Six adjacent slices (each0.25 mm thick) were averaged to generate a low noise point spreadfunction (PSF). Orthogonal slices through the 2-D modulation transferfunction (MTF) were calculated by first computing the Radon transform ofthe point spread function (i.e., integrating along either the x or yaxis), and then calculating the 1-D Fourier transform. Each 1-D profilewas normalized to unity area. A correction was applied to compensate forthe finite diameter of the steel wire. For purposes of comparison, thesame tests were performed on the conventional scanner at 100 kVp for aslice thickness of 1.5 mm. Images were reconstructed using threedifferent reconstruction filters [“High Res Head (#1 H),” “Standard Head(#2),” and “Smooth Abdomen (#3)”].

[0073] The relative imaging performance of the cone beam computerizedtomography system 300 and the conventional scanner were compared usingphantoms and small animals. A simple comparison in soft-tissuedetectability was performed with the phantom shown in FIG. 6(b). Theproximity in CT number between each of the six cylinders makes thisphantom a useful test object for examining contrast sensitivity andsoft-tissue detectability, images were acquired of the phantom with boththe cone beam computerized tomography system 300 and conventionalscanners. Multiple high-resolution

[0074] cone beam computerized tomography slices were averaged to producean equivalent slice thickness to that used on the conventional scanner(1.5 mm). Equivalent exposure (2980 mR) and kVp were used in the twodifferent scans.

[0075] A second test of soft-tissue sensitivity was performed by imaginga laboratory rat that had been euthanized for other purposes, FIG. 6(d).A scanning procedure identical to that described above was used,delivering an in-air, on-axis exposure of 2980 mR at 100 kVp for bothsystems. The resulting 3-D data was reconstructed at voxel sizes of0.25×0.25×0.25 mm³. The subject was also scanned on the conventionalcomputerized tomography scanner at a slice thickness of 1.5 mm. Thisscan delivered the same imaging dose as was delivered by the cone beamcomputerized tomography system 300. For purposes of intercomparison, sixslices from the cone beam computerized tomography data set were averagedto produce a slice thickness equivalent to that of the conventionalscan. The imagers were displayed at comparable window and level to allowcomparison.

[0076] The uniformity of response of the cone beam computerizedtomography scanner shown in shown in FIGS. 7(a)-7(d). Axial and sagittalslices through the cone beam computerized tomography 3-D data set areshown. The images demonstrate a relatively uniform response over theentire field of view of the system. A slight non-uniformity ofapproximately 20 CT numbers (2%) is visible in the histogram equalizedregions of the images. This non-uniformity appears as a combined cuppingand capping artifact. The radial profile (FIG. 7(c)) illustrates thispoint further by comparing to the results obtained from the conventionalscanner (dotted line). An internal check of the reconstruction processusing simulated projection data demonstrates that the non-uniformity isan artifact of the reconstruction process and is dependent upon thechoice of filtering parameters. Apart from the non-uniformity inherentto the reconstruction, the response of the cone beam computerizedtomography system 300 is highly uniform, particularly along thez-dimension.

[0077] In addition to demonstrating uniformity of system response, theimages in FIG. 7 also demonstrate uniform noise characteristics with fewartifacts. This is the case for the full range of exposures studied. Themagnitude and uniformity of the noise is demonstrated in FIGS.8(a)-8(d). The noise varies to a slight degree along the radial axis andto a negligible degree along the vertical axis. A slight dependence onradial position is expected due to the differences in transmissionacross the cylindrical water bath. FIG. 8(c) also presents the measureddependence of noise on exposure [also shown below, in relation to FIG.9(b)]. Overall, the cone beam computerized tomography system 300 iscapable of achieving a noise level of approximately 20 CT numbers for anin-air exposure of 6560 MR at isocenter.

[0078] The noise measured for the cone beam computerized tomographysystem 300 as a function of exposure is shown in the top curve of FIG.9(b). The noise is seen to decrease from −80 units at the lowestexposure examined down to −20 units at the highest. Superimposed is aleast squares fit of the form σ=α+b/{square root}X, where σ is the noisein voxel values, X is the exposure in air at the isocenter, and a and bare constants obtained from the numerical fit. This inverse-square rootdependence upon exposure is consistent with basic noise transfer theoryfor x-ray tomographic reconstructions.

[0079] In order to examine the linearity and accuracy of systemresponse, the CT numbers reported by the cone beam computerizedtomography system 300 for a variety of materials (FIG. 6) were comparedto those reported by the conventional scanner. As shown in FIG. 9(b),the CT numbers of the cone beam computerized tomography system 300 agreewell with those of the conventional scanner. The largest discrepancyover the range of CT numbers was 8 units, with an average discrepancy of5.7. The high coefficient of correlation indicates that, over the rangeexamined, the values reported by the cone beam computerized tomographysystem 300 are proportional to attenuation coefficient.

[0080] The voxel noise of the cone beam computerized tomography system300 and the conventional scanner was compared as a function of exposure,shown in FIG. 9(b). Shown by the open circles and dashed lines are theresults for the conventional scanner using the “High-Res Head (#!H)” and“Standard Head (#2)” reconstruction filters. In each case, the noisedecreases with exposure. An exact comparison between the two systemsrequires that both data sets be reconstructed at equivalent voxel sizeand with the same reconstruction filter. The requirement for equivalentvoxel size was achieved by repeating the noise analysis for the conebeam computerized tomography system 300, with the volume data averagedto give a voxel size equivalent to that of the scanner.

[0081] In order to illustrate the effect of the reconstruction filterupon the voxel Poise, reconstructions were performed with both the“High-Res Head” and “Standard Head” reconstruction filters. The noisefor the cone beam computerized tomography system 300 at equivalent voxelsize is shown by the lower solid curve with a least-squares fitsuperimposed. At equivalent voxel size, it is clear that the cone beamcomputerized tomography system 300 has higher noise at lower exposuresthan the “Standard Head” computerized tomography scanner results.Compared to the “High-Res Head” results for the conventional scanner,however, the cone beam computerized tomography system 300 actuallyprovides lower noise at all but the very highest exposures. Clearly,careful matching of reconstruction filters and reconstruction matrix isrequired to permit exact intercomparison of the two systems.Nonetheless, the results obtained using the cone beam computerizedtomography system 300 are encouraging, since the early prototypeflat-panel detector used in this system is known to exhibit a fairlyhigh level of additive electronics noise, a factor of −5-10 higher thanthat achieved, by more recent electronics designs.

[0082] Results of the noise-power spectrum measurements are summarizedin FIGS. 10(a)-10(c). The 2-D noise-power spectrum in the axial plane(FIG. 10(a)) exhibits a spectral shape typical of systems employingfiltered back-projection reconstruction. The spectral density is reduced(but non-zero) near zero-frequency, increases at mid-frequencies due tothe ramp filter (e.g., peaking around −0.5 mm⁻¹), and declines at higherfrequencies by virtue of the low-pass noise characteristics of thesystem (e.g., 2-D image blur and choice of apodisation window). Slicesof the noise-power spectrum along the u_(x) dimension are shown in FIG.10(b) for various exposure levels. Since the mean signal level is fixedfor each case (i.e., CT#=1000 within the water phantom), the noise-powerspectrum decreases with increasing exposure. Specifically, thenoise-power spectrum appears inversely proportional to exposure in afashion consistent with the form of the numerical fits in FIG. 9(b). Asshown in FIG. 10(c), the noise-power spectrum measured at −1.3 R (in airat isocenter) is −30 mm³ near zero-frequency, increases by a factor of−4 at mid-frequencies, and then descends to about the initial level ofspectral density at the Nyquist frequency.

[0083] Superimposed in FIG. 10(c) are the results measured for theconventional scanner using three reconstruction filters, and tofacilitate intercomparison, noise-power spectrum results for the conebeam computerized tomography system 300 are shown for an equivalentvoxel size. For the #2 and #3 filters, the conventional scanner exhibitsa noise-power spectrum with the characteristic shape described above;however, the high-resolution #1 H filter is seen to significantlyamplify high-frequency noise. The cone beam computerized tomographysystem 300 appears to exhibit low-frequency noise-power spectrumcomparable to the conventional scanner using the #2 and #1 H filters.Given that the choice of reconstruction filter can significantly affectnoise and spatial resolution, and considering the two cases that seemmost closely matched the cone beam computerized tomography system300—even in its initial, un-optimized configuration—appears to providenoise performance comparable to the conventional scanner. As evident inFIG. 9(b), the cone beam computerized tomography system 300 exhibitslower voxel noise than the conventional scanner (#1 H) at low exposures.Similarly, the cone beam computerized tomography system 300 exhibitsreduced high-frequency noise-power spectrum. These initial results areespecially promising considering the on-going improvements in FPI designand readout electronics.

[0084] The response of the cone beam computerized tomography system 300to the wire test object is presented in FIG. 11(a). Overall, the PSF issymmetric (aside from a small streak artifact believed associated withthe image lag characteristics of the system) and has a full-width athalf-maximum (FWHM) of 0.6 mm. The system MTF is shown in FIG. 11(b) forboth the on- and off-axis wire results. These results suggest that thefrequency pass of the system in the z=0 plane does not changesignificantly over the relatively s mall (−30 mm) range examined. Thestrong influence of the reconstruction filter is demonstrated in the MTFresults for the conventional scanner, also shown in FIG. 11(b).

[0085] The “Standard Head (#2)” filter significantly reduces the signalpass of the system compared to the High-Res Head (#1 H)” filter. Theresults demonstrate that the MTF of the conventional scanner iscomparable to that of the cone beam computerized tomography system 300when the “High-Res Head (#1 H)” filter is used. This observation isconsistent with the noise results presented in FIG. 9(b). The resolutionof the cone beam computerized tomography system 300 and conventionalscanner have not been compared in the z-dimension. It is expected,however, that the spatial resolution of the cone beam computerizedtomography system 300 in the z-dimension will be comparable to thatmeasured in the axial plane. Of course, the spatial resolution of theconventional scanner will be limited by the selected slice thickness,which is typically 1 mm or greater. The nearly isotropic resolution ofthe cone beam computerized tomography system 300 is expected to be asignificant advantage for detection and localization.

[0086] FIGS. 12(a) and 12(b) show axial image slices of the low-contractphantom obtained on the cone beam computerized tomography system 300 andthe conventional computerized scanner at equivalent kVp and exposure.The grayscale window in each case is quite narrow in order to maximizethe displayed contrast, and despite the slight signal non-uniformityevident for the cone beam computerized tomography image (cupping/cappingartifact discussed above) the visibility of each insert is comparable tothe conventional scanner. The mean signal values for each material areas shown in FIG. 9(a). Slight differences in system response (e.g., dueto detector response, x-ray spectrum, etc.) can result in contractreversal for materials with CT# very close to that of water. For examplein the case of the brain insert (lower right), even the slight (−5 CT#)difference between the mean value reported by the cone beam computerizedtomography system 300 and the conventional scanner is sufficient to givean apparent inversion in the density of the material relative to water.The minimum detectable contrast is arguably superior for the cone beamcomputerized 20 tomography system 300 (e.g., visibility of the brain andCB-3 inserts), but this remains to be verified by a more controlled,quantitative observer study.

[0087] The overall performance of the cone beam computerized tomographysystem 300 is demonstrated in the images of the volumetric data setillustrated in FIGS. 13(a)-13(i). These images of an euthanized ratdemonstrate the soft tissue sensitivity and high spatial resolution ofthe system. Example images are shown from various regions throughout thevolumetric set [e.g., in regions of the lungs (a, b, c), the kidney (d,e, f), and lower spine (g, h, i)] to illustrate the quantity and uniformquality of the data produced with the cone beam computerized tomographysystem 300. The clear visualization of soft-tissue structuresdemonstrates the soft-tissue contrast sensitivity of the scanner.

[0088] In FIGS. 13(a)-13(c), the window and level have been set toemphasize features in the lung of the rat. In addition to the lungdetail, there are some streak artifacts evident, the origin of which isunknown, but is believed to be associated with detector lag effects orbeam hardening.

[0089] The soft tissue contrast sensitivity of the cone beamcomputerized tomography system 300 is illustrated in FIGS. 13(d)-13(f),in which the window and level have been set to delineate fat and muscle.The cross-hair in each image indicates the location of the rat's leftkidney. These images illustrate the advantage of a nearly isotropicspatial resolution for delineation of a 3-D structure such as thekidney. Other structures, such as the stomach, bowel and liver are alsoclearly visible.

[0090] The spatial resolution performance of the system 300 isdemonstrated in FIGS. 13(g-i), in which the same rat data set isdisplayed with window and level selected to display bony features. Theclear visibility of the intervertebral spaces and the non-cortical bonein the pelvis is stunning. It should be kept in mind that this level ofdetail was produced on a cone beam computerized tomography system 300that operates on a scale that mimics the geometry of the linearaccelerator. Therefore, this level of detail would be expected in theclinical implementation of the device, given accurate correction ofmechanical flex. The volumetric data set is illustrated further in FIG.14, in which volume renderings demonstrate the fully 3-D nature of thedata set and show the level of detail contained within the cone beamcomputerized tomography data. It is interesting to note that all thedata presented in FIGS. 13 and 14 were obtained from a singleacquisition performed in a single rotation.

[0091] Finally, the quality of images produced by the cone beamcomputerized tomography system 300 was assessed by comparison to imagesproduced by the conventional scanner. FIGS. 15(a)-15(b) show an axialslice of the rat acquired on the two systems. At equivalent exposure,the images produced by the two systems are of comparable quality both interms of spatial resolution and contrast sensitivity. The flat panelimager-based cone beam computerized tomography image exhibits exquisitespatial resolution and provides clear delineation of soft-tissueboundaries and detail in the gut. The spatial resolution of the conebeam computerized tomography system 300 appears to exceed that of theconventional scanner; however, it must be noted that restrictions inavailable reconstruction matrices for the conventional computerizedtomography scanner limited the voxel size to twice that of the cone beamcomputerized tomography image. Lack of obvious pixelation in the flatpanel imager-based cone beam computerized tomography image indicatesthat this level of detail represents the physical limits in spatialresolution of the current system.

[0092] The objective of these investigations is to evaluate theapplicability of flat-panel technology as a detector in a cone beamcomputerized tomography system, specifically, a tomographic imagingsystem for use in the guidance of radiation therapy on a medical linearaccelerator.

[0093] The quantitative and qualitative results of our studies suggestthat a cone beam computerized tomography scanner based on flat paneldetector technology is a viable means for high performance computedtomography. Initial studies of signal response uniformity demonstratedthat the response of the system is uniform over the field of view towithin ±2%, with the slight degree of non-uniformity apparent as acombined cupping and capping artifact in the x-y plane attributable to areconstruction artifact. The linearity of response was demonstratedusing a range of soft-tissue test materials and was found to be linearto within ±−0.6%. Measurements of image noise versus exposuredemonstrate that the prototype cone beam computerized tomography system300 performs comparably to the conventional scanner, demonstrating theinverse square root exposure dependence predicted by theory.Investigations of noise power spectrum and spatial frequency responsefor the two systems reinforce these conclusions and illustrate theadvantages of developing more extensive (empirical and theoretical)frequency-dependent characterization methods for volumetric computedtomography systems.

[0094] In addition to the quantitative measures of performance, theimages of low-contract phantoms and small animal anatomy confirm theconclusions drawn from these measures, showing excellent detail andsoft-tissue contract, more than sufficient for tissue localization inradiation oncology.

[0095] The results presented here demonstrate the potential of thisapproach for volumetric imaging. However, this study has been performedunder conditions of small object size and small cone angle. Theseconditions are imposed by the size of the detector used in thisinvestigation. Imaging with larger detectors allows increased cone angleand, for computerized tomography, increased object thickness. Theextrapolation of performance based on the results presented here to thatfor larger detectors must be done with some caution. Imaging largerobjects with an increased field of view will result in increased scatterand reduced transmission. The increase in scatter can be expected tohave a negative impact on computerized tomography imaging performance byintroducing non-uniformities in the reconstructed image (e.g., cuppingand/or streaks), and by adding additional x-ray quantum noise to theimage signal. The magnitude of scatter reaching the detector will dependgreatly on the cone-angle and air gap employed, and studies suggest thatscatter at these distances may be reduced compared to conventionalradiographic applications. Quantifying the magnitude of the x-rayscatter problem and developing methods to reduce it are areas of ongoinginvestigation.

[0096] In addition to concerns of x-ray scatter at large cone-angles,the scanning of larger objects will significantly reduce the fluencearriving at the detector. This reduced transmission will negativelyimpact the performance of the flat-panel detector. Currently availableflat panel imagers demonstrate performance inferior to conventionalimage intensifiers at fluoroscopic exposure rates, due to the presenceof additive noise in the flat-panel readout electronics. Additive noisecauses the detected quantum efficiency of the imager to depend on thenumber of x-rays forming an image. This dependence is illustrated inFIG. 16 for the flat-panel imager 306 used in these investigations andfor hypothetical detectors that embody the most recent advances inimager 306 design, including higher x-ray quantum detection efficiencythrough the use of Csl:TI and a reduction in additive noise throughimprovements in readout electronics.

[0097] The zero-frequency detected quantum efficiency was computed usinga model for signal and noise transfer that has demonstrated excellentagreement with measurements. It is clear from FIG. 16 that improvementsin the x-ray converter and electronics noise significantly reduce theexposure dependency of the detected quantum efficiency over the broadrange of exposures required for computerized tomography. The magnitudeof the reduction depends greatly on the amplifier noise in the system.For the prototype imager used in these studies, the amplifier noise isvery high at 12,700 e. For the low transmitted exposure levels incomputerized tomography of pelvic anatomy, for example, this detectorwould achieve a zero-frequency detected quantum efficiency of less than10%. In comparison, an imager than incorporates the recent advances indesign listed above (e.g., a high-quality Csl:TI converter and amplifiernoise of 3000 3 or better) would achieve a higher detected quantumefficiency (−65%) at full transmission and maintain a detected quantumefficiency of>40% even at the low exposure levels. Such enhancements inimager design are within the current capabilities of flat panel imagermanufacturers and will greatly facilitate the application of flat panelimagers in cone-beam computerized tomography of human beings.Furthermore, these improvements are largely driven by other forces indigital imaging that anticipates use of flat panel imagers in place ofconventional image-intensifier systems for interventional fluoroscopy.For this reason, it can be expected that imagers with suchcharacteristics will be available within the next five years.

[0098] Overall, the operating characteristics of the flat-panel arehighly compatible with acquisition in a cone beam computerizedtomography scanning geometry. Unlike image-intensifier or lens basedsystems, flat panel detectors are geometrically robust under a rotatinggeometry, eliminating concerns of image distortion. The proximity of theanalog-to-digital converter to the pixel element and the relativelylarge charge signals make the panels robust in high radio-frequencypower environments; this is of particular interest for radiotherapyapplications. The high readout rate of these detectors allows forimaging sequences of 300 projection images to be acquired within 10seconds (operating at 30 fps). This is more than sufficient to satisfythe allowable rotation rates for the gantry of a medical linearaccelerator. In fact, while the International ElectromechanicalCommission (IEC) recommends less than 1 revolution per minute for linearaccelerators, it would be reasonable to reconsider such constraints inlight of the advantages of cone beam computerized tomography guidance inthe treatment room. Currently, the detector size and aspect ratio aredriven by the needs of digital radiography, producing detectorscomparable in size to radiographic film. These sizes limit thefield-of-view of the reconstruction if sufficient clearance is to bemaintained between ft detector and patient during gantry rotation. Thisproblem can be addressed using offset detector schemes that use 360° ofgantry rotation. Ultimately, a specialized detector could be designedwith a size and aspect ratio that match the requirements for cone beamcomputerized tomography (e.g., a −25×50 cm² area panel).

[0099] Given the potential that this technology is demonstrating, theopportunities for new areas of application for computed tomography aresignificant. Imaging systems based on this technology can be constructedto address specific imaging problems, including non-destructive testing(at kilovoltage or megavoltage energies), early detection and monitoringof specific medical conditions, and, of course, navigational imaging fortherapies. The compact nature of the panels allow flat panelimager-based cone beam computerized tomography imagers to be applied insituations that would never be considered feasible for a conventionalcomputerized tomography scanner. The cone beam computerized tomographyapproach offers two important features that dramatically reduce its costin comparison to a conventional scanner. First, the cone-beam nature ofthe acquisition does not require an additional mechanism to move thepatient (or object) during image acquisition. Second, the use of acone-beam, as opposed to a fan-beam, significantly increases the x-rayutilization, lowering the x-ray tube heat capacity required forvolumetric scanning. For the same source and detector geometry, theefficiency roughly scales with the slice thickness. For example, thex-ray utilization increased by a factor of 30 in going from a 3 mm slicein a conventional scanner to a cone-angle corresponding to a 100 mmslice with a cone-beam system, This would decrease heat-load capacitiesdramatically. From our experience, a 5200 kHU x-ray tube costsapproximately $70,000, whereas a 600 kHU x-ray tube (a factor of −10lower in capacity) costs roughly $6,000.

[0100] Cone-beam computed tomography has been a topic of active researchand development for over a decade in areas such as nuclear medicine andindustrial testing; however, only recently has it begun to appear in thediagnostic computerized tomography arena. The developments in this areahave been for the most part limited to multi-slice detectors. In thisinvestigation, the use of an alternative detector for high-qualitycomputerized tomography has been studied. The results of theinvestigation suggest that there is a significant potential for the useof these detectors in cone beam computerized tomography systems forradiotherapy and quite possibly for diagnostic and interventionalcomputerized tomography imaging tasks that will take advantage of thefully 3-D nature of cone beam computerized tomography.

[0101] Based upon the positive results presented previously with respectto the cone beam computerized tomography system 300, several embodimentsof a flat panel imager-based kilovoltage cone beam computerizedtomography scanner for guiding radiation therapy on a medical linearaccelerator are envisioned. For example, FIGS. 17(a)-(e) and 18 arediagrammatic and schematic views of an embodiment of a wall-mounted conebeam computerized tomography system 400. The cone beam computerizedtomography system 400 includes an x-ray source, such as x-ray tube 402,and a flat-panel imager 404 mounted on a gantry 406. The x-ray tube 402generates a beam of x-rays 407 in the form of a cone or pyramid thathave an energy ranging from approximately 30 KeV to 150 KeV, preferablyapproximately 100 KeV. The flat-panel imager 404 employs amorphoussilicon detectors.

[0102] The system 400 may be retrofitted onto an existing or newradiation therapy system 700 that includes a separate radiation therapyx-ray source, such as a linear source 409, that operates at a powerlevel higher than that of x-ray tube 402 so as to allow for treatment ofa target volume in a patient. The linear source 409 generates a beam ofx-rays or particles 411, such as photons or electrons, that have anenergy ranging from 4 MeV to 25 MeV. The system 400 may also include animager (not shown) that is aligned with the linear source 409 with thepatient interposed therebetween. The imager forms projection images ofthe patient based on the remnants of the beam 411 that passes throughthe patient. Note that the x-ray sources 402 and 409 may be separate andcontained with the same structure or be combined into a single sourcethat can generate x-rays of different energies.

[0103] As shown in FIGS. 17(a)-(e) and 18-19, the flat-panel imager 404can be mounted to the face of a flat, circular, rotatable drum 408 ofthe gantry 406 of a medical linear accelerator 409, where the x-ray beam407 produced by the x-ray tube 402 is approximately orthogonal to thetreatment beam 411 produced by the radiation therapy source 409.Attachment of the flat plane imager 404 is accomplished by an imagersupport system 413 that includes three 1 m long arms 410, 412 and 415that form a tripod. Side arms 410 and 415 are identical to one anotherin shape and have ends attached to a Ax95 Guy pivot 417 which in turn isattached to a mounting 414 by screws that are threaded through alignedthreaded holes of the pivot 417 and threaded holes 425 and 431 of plates433 and 435, respectively, as shown in FIGS. 18 and 19(a)-(b). As shownin FIGS. 17(b) and 18, the mountings 414 for the arms 410 and 415 arealigned with one another along a line segment 419 that is containedwithin a plane 421 that is parallel to and offset by approximately 30 cmfrom the plane containing the flat-plane imager 404. The mountings 414are separated from one another by approximately 70 cm and aresymmetrically positioned with respect to a plane bisecting an imagermount 423 that is attached to the drum 408 270° from the radiationtherapy source 409.

[0104] As shown in FIGS. 18 and 19(a)-(b), each mounting 414 is attachedto an end portion 416 of the drum 408 by inserting a threaded malemember 418 through an opening 437 formed through the drum 408. Onceinserted, the male member 418 is attached to the drum 408 by tighteninga nut 420 onto the threaded male member 418. The other ends of the arms410 and 415 are attached to Ax95 Guy pivots 422 attached to the back ofan ⅜ inch thick Aluminum square plate 424 is attached to the rear of theflat-panel imager 404 via bolts (not shown).

[0105] As shown in FIGS. 17(d)-(e), there are two preset positions ofthe flat panel imager 404 relative to the plate 424. As shown in FIG.17(d), the flat panel imager 404 is centered about the ends of the arm412. In order to provide a larger field of view, an offset flat panelimager 404 can be used as shown in FIG. 17(e) where the imager 404 isattached to a side of the plate 424 via bolts. Note that it is possibleto use a motorized system to move the flat panel imager 404 relative tothe plate 424 to provide an easy way to vary the field of view of a conebeam computerized tomography system.

[0106] A center arm 412 is also attached to the drum 408 and theflat-panel imager 404. The center arm 412 has one end attached to Ax95Guy pivot 427 that is in turn attached to a tapped, triangular-shaped,reinforcing plate 426 formed on the drum 408 as shown in FIGS. 17(b) and18. The plate 426 is approximately 433.8 mm from the rotational axis 428that intersects the iso-center 430 of the imaging system 400. A secondend of the center arm 412 is attached to the plate 424 via a Cx95A rightangle joint 425.

[0107] As shown in FIGS. 17(b) and 18, the end of the arm 412 lies alonga line that is the perpendicular bisector of the line segment 419 and isradially separated from the midpoint between mountings 414 as measuredalong line segment 419 by a distance D of approximately 30 cm.

[0108] As shown in FIGS. 17(b) and 18, the other ends of the arms 410,412 and 415 are attached to the plate 424 so as to be positionedapproximately 20 cm from the rear edge 429 of the plate 434 andapproximately midway between the left and right edges of the plate 434.

[0109] Once the arms 410, 412 and 415 are attached to the drum 408 andthe plate 424, the arms can be pivoted so that the flat panel imager 404moves to a position where its rear side is separated from the iso-center430 by a distance L of approximately 600 mm. One advantage of the imagersupport system 413 is that it can be used to retrofit existingstand-alone radiation treatment devices so they have the capability tohave a flat panel imager attached thereto. The imager support system 413is very rigid, i.e., constant tension and compression, which reducesmovement of the imager 404 and so leads to cleaner imaging data.

[0110] Note that the x-ray tube 402 can also be retrofitted onto anexisting stand-alone treatment device so as to be positioned opposite tothe flat panel imager 404. As shown in FIGS. 17(a)-(e), the x-ray tube402 is attached to tube support 440 that is composed of a pair of frontand rear faces 442 and 444 and a pair of side faces 446. A multi-leafcollimator 448 is supported within the interior of the tube support 440.The front and rear faces 442 and 444 each include three openings 450,452 that are aligned with one another and receive three cylindricalsupport arms 454 that are attached to a bearing housing 456 that isbolted to the drum 408. The tube support 440 and the x-ray tube 402 areable to slide along the support arms 454. Note that a cable support 458spans between the tube support 440 and the bearing housing 456 andcontains the wiring necessary to operate the x-ray tube 402.

[0111] An alternative imager support system for the flat panel imager404 of FIGS. 17 is shown in FIGS. 20(a)-(b). In particular, the imagersupport system 507 shown in FIGS. 20(a)-(b) includes a single pivotingarm 510 that has one end 511 pivotably attached to a lower corner of theradiation therapy source 409. The other end 512 of the arm 510 ispivotably attached to an end of the flat-panel imager 404. The arm 510and flat-panel imager 404 are movable from a retracted position of FIG.20(a) to an extended position of FIG. 20(b) and vice versa. Movement ofthe arm 510 and the flat-panel imager 404 may be moved either manuallyor via a motor.

[0112] Note that when the imager support system 507 is used, the x-raytube 402 is attached to a second lower corner of the radiation therapysource 409 in order to simplify the support structure and reduce themechanical complexity of the overall system. The position of the x-raytube 402 also reduces interference with staff access to the patient.Note that in this embodiment, the distance from the x-ray tube 402 tothe axis of rotation 428 is not necessarily equal to the distance fromthe radiation therapy source 409 to the axis of rotation 428. Also, theamount of extension of the arm 510 shown in FIG. 20(b) will varydepending on the desired field of view for cone beam computerizedtomography imaging. Note that if the mechanics are engineered to besufficiently precise, the arm 510 can move in and out during imageacquisition during gantry rotation so as to allow the imager 404 todynamically avoid potential rotation-induced collisions with the patientor the table. The head of the radiation therapy source 409 can bealtered to provide additional lead shielding on the imager side to limitradiation induced damage to the imager 404 while in the retractedposition of FIG. 20(a). This will increase the life span of the imager404.

[0113] A second alternative imager support system for the flat panelimager 404 of FIG. 17 is shown in FIGS. 21 (a)-(b). In particular, theimager support system 607 shown in FIGS. 21 (a)-(b) includes a singleC-arm 610 that is attached to an arm support 611 that is attached to thefront or rear of the radiation therapy source 409. At one end of theC-arm 610 the x-ray tube 402 is attached and at the other end theflat-panel imager 404 is attached. The C-arm 610 is moved eithermanually or by a motor within the arm support 611, so that the x-raytube 402 and the flat-panel imager 404 can move along an arc.

[0114] Note that in this embodiment, the distance from the x-ray tube402 to the axis of rotation 428 is not necessarily equal to the distancefrom the radiation therapy source 409 to the axis of rotation 428. Thearm 610 does not necessarily be in the shape of an arc of a circle.Also, the axis of rotation of the arm 610 is not necessarily coincidentwith the axis of rotation 428 of the radiation therapy source 409, whichallows the same device to be fitted on machines with differentface-to-isocenter distances without altering the radius of the C-arm610.

[0115] Use of the C-arm 610 of FIGS. 21 (a)-(b) allows for a greatamount of flexibility in obtaining cone beam computerized tomographyimages. For instance, image data can be obtained by only having the drum408 of the gantry 406 rotate. Image data can be obtained in a secondmanner by having the C-arm 610 move independently of the gantry 406 in acircular path. Image data can be obtained by having the C-arm 610 andthe drum 408 work cooperatively to generate images along a circular pathso that the angular range of acquisition is increased and soinstabilities in the angular velocity of the gantry are addressed. Afourth manner of imaging involves rotating the drum 408 and pivoting theC-arm 610 about the mounting point on the gantry 406 with a sinusoidalpattern to effect non-circular orbits that involve a sinusoidaltrajectory on a spherical surface. Such a non-circular orbit allows morecomplete image reconstructions by satisfying Tuy's condition.

[0116]FIG. 22 shows a portable cone beam computerized tomography system700. In this embodiment, the system 700 is on a mobile platform 702 sothat it can be moved relative to a patient 441 positioned on a table 443relative to a rotating radiation therapy source 409 (not shown). Thecone beam computerized tomography system 700 includes an x-ray source,such as x-ray tube 402 positioned on one side of a C-arm 704, and aflat-panel imager 404 positioned on an opposite side of the C-arm 704.The C-arm 704 can rotate about two axes of rotation when in operation.The system 700 can be moved to a radiation therapy system (not shown)and can be used to generate images that aid in the alignment of theradiation therapy system.

[0117] With the above descriptions of the cone beam computerizedtomography system 400 and the various embodiments of the imager supportsystems shown in FIGS. 18-22 in mind, operation of the system 400 isdescribed below. In the description to follow, the term “shape” of theradiation therapy beam 411 is understood to refer to the spatialdistribution of the beam in a plane perpendicular to the direction ofthe beam or to the frequency modulation of the beam after beingtransmitted through some beam-limiting device. The term “planning image”refers to an image of the patient acquired by the cone beam computerizedtomography system 400 prior to treatment delivery used for radiationtherapy treatment planning. The term “constrained plan set” refers to aplurality of radiation therapy treatment plans for a given patient,where each radiation therapy treatment plan is calculated assuming someperturbation of lesion location and/or orientation compared to that inthe planning image. For example, a constrained plan set could becalculated where each plan corresponds to a different magnitude oflesion rotation about the y and/or z axes.

[0118] The cone beam computerized tomography imaging system 400preferably includes an x-ray tube 402 and a flat panel imager 404 in anyone of the geometries illustrated in FIGS. 23(a)-(d), capable of forming3-D images of the patient on the treatment table in the treatmentposition. The x-ray tube 402 may be operated so as to produce a pulsedor continuous beam of x-rays 407. The flat panel imager 404 includes anactive matrix of imaging pixels incorporating mechanisms for: 1.)converting incident x-rays to electronic charge (e.g., a scintillator incombination with optically sensitive elements at each pixel, or aphotoconductor); 2.) integrating and storing the electronic charge ateach pixel (e.g., the capacitance of photodiode(s), capacitors, etc.located at each pixel); and 3.) reading the electronic charge out of thedevice (e.g., a thin-film transistor switch or the like at each pixel,with associated switching control lines and readout lines). The x-raytube 402 and the flat panel imager 404 preferably move in a circularorbit (or variation thereof) about the longitudinal axis of the patient.Depending on which ones of the imager support systems used in FIGS.18-22, the imager support system should accommodate offsets in the xand/or z directions as illustrated in FIG. 23(b). Note that the combinedmotion of the x-ray tube 402 and/or the flat panel imager 404 in x, y,and/or z is termed the orbit, and may be circular about the patient, ornon-circular, e.g., comprising of some combination of linear,sinusoidal, circular, and/or random paths. For example, in the casewhere the source 402 and imager 404 move independently with respect toone another, the source 402 can move on a sinusoidal or sawtooth pathconstrained to the surface of a cylinder while the imager 404 moves in acircular path on the surface of acylinder. In this scenario, thecollimator adjusts in real time the shape of the radiation field so itis confined to the imager 404 despite the allowed independent motion ofthe source 402 and imager 404.

[0119] Cone beam computerized tomography image acquisition involvesacquisition of a plurality of 2-D images, where each image preferablycorresponds to a different orientation of the x-ray beam 407 and theflat panel imager 404 with respect to the patient 441, e.g., where thex-ray tube 402 and the flat panel imager 404 traverse a circular ornon-circular path about the patient 441 as illustrated in FIG. 23(d).Note that the cone beam computerized tomography image is preferablyacquired with the patient on the treatment table, in the treatmentposition, and immediately prior to treatment delivery. The processesinvolved in the preferred method for cone beam computerized tomographyimage acquisition are illustrated in FIG. 24, divided conceptually intoa variety of off-line and on-line processes, and mechanisms for 2-Dimage acquisition and 3-D image reconstruction.

[0120] The off-line processes schematically shown in FIG. 24 includeacquisition of a plurality of 2-D images acquired in the absence ofx-ray irradiation (termed dark fields) and with uniform x-rayirradiation (termed flood fields). Such dark and flood fields are usedto correct stationary nonuniformities in the imaging system arising fromnonuniformity in pixel operational and response characteristics. Alsoincluded is a mechanism for identifying and correcting defective pixelsin the 2-D images (e.g., a pixel defect map that identifies aberrantpixel coordinates, and application of a filter to the correspondingpixel values). Thirdly, a measure and process for correction of orbitnon-idealities, described below, is preferably employed.

[0121] The on-line processes schematically shown in FIG. 24 include: 1.)control and monitoring of the x-ray tube; 2.) control and monitoring ofthe orbit traversed by the x-ray tube 402 and the flat panel imager 404(e.g., by rotating the gantry 406); and 3.) control and readout of theflat panel imager 404. The x-ray source 402 produces x-rays in either apulsed or continuous manner, and variations in the magnitude of x-raytube output are monitored by an output monitor, which preferablyincludes a radiation sensitive electronic device such as a diode placedinside the x-ray tube collimator assembly. Alternatively, the outputmonitor could be placed outside the x-ray tube 402 in a position thatallows it to measure variations in x-ray tube output, or the outputcould be measured using pixels on the flat panel imager 404, such thatthese pixels are not occluded by the patient in the plurality of 2-Dprojection images. The orbit of the x-ray tube 402 and the flat panelimager 404 about the patient is preferably controlled viacomputer-controlled rotation of the gantry 406, combined with a precisemeasurement of the gantry angle at which each 2-D image is acquired. Forembodiments in which the x-ray source 402 and the flat panel imager 404are not both mounted on the treatment gantry 406, such as the portableembodiment of FIG. 22, a similar mechanism for measuring and recordingthe location of these two components for each 2-D image is employed.Thirdly, a plurality of 2-D images are read from the flat panel imager404 by a controvacquisition computer. The readout of the flat panelimager 404 is preferably synchronized with the operation of the x-raytube 402 as well as with the rotation of the x-ray tube 402 and the flatpanel imager 404 support structure(s), such as those describedpreviously with respect to FIGS. 18-22. The timing of x-ray exposures,gantry rotation, and flat panel imager readout is preferablysynchronized by: 1.) the control/acquisition computer; or 2.) anexternal trigger mechanism (gating source), such as a device for activebreathing control, electrocardiac gating, etc. For the former case, thepreferred embodiment includes computer-control of: 1.) x-ray pulsesgenerated by the x-ray source 402; 2.) gantry rotation (e.g., inincrements of ˜1° through ˜360°); and flat panel imager readout (e.g.,at a readout rate consistent with the limitations in x-ray tube outputand gantry rotation). For the latter case, the preferred embodiment issuch that the gating source triggers x-ray production, gantry rotation,and flat panel imager readout in a manner synchronized with the motionof anatomical structures in the patient 441 in order to reduce thedeleterious effects of organ motion in image reconstructions.

[0122] The preferred embodiment includes a mechanism (reconstructionengine) for high-speed cone beam computerized tomography imagereconstruction. The plurality of 2-D projections is first processed bydark and flood field correction, and the measurements of orbitnon-ideality (below), tube output variations, and gantry rotation areused together with the processed 2-D projections to form 3-D cone beamcomputerized tomography image reconstructions of the patient 441. Avariety of cone-beam reconstruction techniques are known within the art,including cone-beam filtered back-projection. The cone beam computerizedtomography image is then made available to a system for on-linetreatment planning.

[0123] In the interim between the 2-D image acquisition and correctionof lesion localization errors, the patient 441 is preferably monitoredby periodic radiographs obtained with the flat panel imager at one ormore gantry angles. In the preferred embodiment, these monitorradiographs are analyzed (e.g., by calculation of difference images) inorder to provide a check against intrafraction motion of the patient441.

[0124] The preferred embodiment includes a computer-controlled treatmenttable 443 for correction of lesion localization errors. The table 443preferably allows translation of the patient 441 in the x, y, and zdirections as well as rotation about the x axis. Rotation about the yaxis (tilt) and z axis (roll) is possible for an embodiment in whichlesion localization errors are corrected by such motions (as opposed tocorrection of such errors through selection of an appropriate RTTP froma constrained plan set), provided that such motions do not causeuncertainty in the location/orientation of the lesion 444 and/orsurrounding structures, e.g., due to the effects of gravity.Furthermore, the treatment table 443 is preferably constructed ofradio-translucent material so as not to interfere significantly with theacquisition of cone beam computerized tomography images.

[0125] The preferred embodiment includes a method for calibration of theradiation therapy delivery system accomplished using a radiation therapysystem including the radiation therapy source 409, a collimatingstructure such as a multi-leaf collimator, and an imager 446. The imager446 is located opposite the radiation therapy source 409 on a supportarm attached to the radiotherapy gantry 406 and in the preferredembodiment is a flat panel imager 404 designed for imaging of the highenergy beam 411. The calibration method preferably employs a referenceBB 448 embedded in a lucite cube 450 and located at a known locationwith respect to the isocenter 430 of gantry rotation, as illustrated inFIG. 25. The cube 450 is precisely leveled, and marks on the cubesurface project the location of the isocenter within the cube. Thereference BB 448 is imaged at angular increments using the radiationtherapy source 409 and imager 446 as the gantry 406 rotates through360°, preferably clockwise and counter-clockwise. In each image, thereference BB 448 is located preferably by an automated centroidcalculation, and the edge of each leaf of the multi-leaf collimator andthe edge of the collimators are localized by calculation of maximumsignal gradient. After subtracting a sinusoid of specified amplitudefrom the measured deflections, the residuals represent imperfections inleaf placement. These residuals can then be incorporated into thecontroller of the multi-leaf collimator and calibrated out. Analternative approach is to modify the planning system to generate“corrected” leaf positions. After calibration, the patient positioninglasers in the treatment room are adjusted to the set of laser alignmentmarks located on the lucite cube.

[0126] The preferred embodiment furthermore includes a calibrator thatcalibrates the cone beam computerized tomography imaging geometryrelative to that of the radiation therapy source 409. Calibration of thecone beam calibration tomography geometry is preferably performedimmediately following multi-leaf collimator leaf calibration, withoutmoving the reference BB 448. The same procedure is performed using thex-ray source 402 and the flat panel imager 404;

[0127] however, in this case, the residuals are used to adjust theback-projection trajectories in the reconstruction process. The image ofthe localized BB 448 is preferably analyzed using a 3-D centroidalgorithm, and the location of isocenter 430 is calculated as a simpleoffset from the centroid. The isocenter 430 can thus be explicitlyidentified within the 3-D matrix of cone beam computerized tomographyimages.

[0128] In the preferred embodiment, the cross-calibration of the conebeam computerized tomography imaging system 400 and the radiationtherapy delivery system can be tested with a mechanism (phantom) forcombined geometry and dosimetry measurement. The phantom preferablyincludes a water-filled or water-equivalent volume in which a dosimetryinsert is rigidly placed at various locations. The dosimetry insertpreferably contains either: 1.) a detector matrix of electronicdosimeters, or 2.) a volume of radiosensitive gel dosimeter. In theformer case, the dosimeters are embedded in a water-equivalent insertand placed asymmetrically to allow unambiguous identification in acomputerized tomography image; furthermore, each dosimeter issufficiently small as to have legible influence on the dosimetry ofother detectors. The electronic signals from the dosimeter matrix arepreferably used in either of two ways: 1.) the dosimetry of a completedelivery can be tested by recording the signal from all detectors andcomparing to calculations, thereby providing a point dose verificationof the delivery as well as routine pretreatment quality assurance;and/or 2.) the precision and accuracy of the combined imaging anddelivery system can be measured by recording the dose to the detectorsas the geometric edge of a leaf can be inferred and compared to theplanning system dose calculation. This test is preferably performed forall the leaves in the system by moving the location of the dosimetryinsert within the volume. In the case of a radiosensitive gel dosimeter,measurement of 3-D dose distributions delivered by a given treatmentscheme can be quantitatively evaluated.

[0129] The preferred embodiment furthermore includes delineating thetarget volume immediately following acquisition of the cone beamcomputerized tomography image of the patient 441 on the treatment table443 in the treatment position. Localization of the target volume/lesion444 and/or surrounding structures can be performed manually, e.g., bycontouring of structures in some combination of transaxial, sagittal,coronal, and/or oblique slices. Alternatively, the target volume/lesion444 and/or surrounding structures can be delineated by an automatedlocalization algorithm, as recognized in the art. In this approach, thetarget volume/lesion 444 defined in the planning image is overlaid on agiven on-line cone beam computerized tomography image, and the imagesare matched, e.g., by translating and rotating the reference targetcontour in such a way as to minimize the standard deviation of pixelvalues compared to the planning image. In the planning image, bonystructures are defined, and matching of the planning image with theon-line cone beam computerized tomography image (both with calibratedisocenter positions) on bony structures determines the setup error(rotation and translation) of the bony anatomy. The motion of thesoft-tissue target relative to the bony anatomy is quantified bytranslating and rotating the target volume contours until they cover ahomogeneous area (i.e., standard deviation in pixel value differences isminimized).

[0130] The treatment plan for the current session can be modified basedon the cone beam computerized tomography image data by a number ofmethods or combinations therein, including recalculation of the RTTP,selection of a modified RTTP from a previously calculated set of plans,and/or translation, rotation, and/or angulation of the patient. Themethod chosen should provide a modified plan for the current treatmentsession in a manner that does not cause uncertainty in thelocation/orientation of the lesion; therefore, the method should becompleted within a short time frame in order to minimize intrafractionorgan motion effects, and should not significantly distort patientanatomy. Recalculation of the RTTP based on the cone beam computerizedtomography image data should be consistent with such time constraints.Similarly translation, rotation, and/or angulation of the patient shouldnot perturb patient anatomy compared to that measured in the cone beamcomputerized tomography image, e.g., due to the effects of gravity.

[0131] The preferred embodiment entails a streamlined process for rapidlesion localization, selection of an appropriate RTTP, dosimetry review,and transfer of the prescription to the radiation therapy deliverysystem. The process for on-line cone beam computerized tomographyguidance of radiation therapy procedures is illustrated in FIG. 26,which conceptually separates the system into: 1.) the off-line treatmentprocess; 2.) priors for on-line selection and correction; and 3.) theon-line imaging and treatment process.

[0132] The off-line treatment process in the preferred embodiment beginswith a planning image on which contours of the target volume andsurrounding structures are defined, and margins for target deformation,delivery precision, and delineation precision are applied. Inverseplanning is performed according to a given protocol for radiationtherapy of the given treatment site, e.g., a number of radiation therapybeams 411 directed at the patient 441 from various angles, with targetdose uniformity and normal tissue volume constraints to match theprescription. In addition to this reference plan, a plurality ofadditional plans (the constrained plan set) are generated as a functionof various translations and/or rotations of the target volume. Plans arepreferably generated at small increments of each possible translationand/or rotation (e.g., rotation of the target volume about the y axis).

[0133] In the preferred embodiment for on-line plan selection andcorrection of lesion localization errors, the target volume/lesion 444and its relationship to bony structure in the planning image areprepared for use as priors, and the constrained plan set is transferredto the radiation therapy system to verify deliverability prior to theon-line procedure. In the on-line treatment process, the patient 441 isset up on the treatment table 443 in the treatment position, and conebeam computerized tomography images are acquired as described above. Thetarget volume/lesion 444 and surrounding structures are delineated inthe cone beam computerized tomography data, thereby identifying thetranslations and/or rotations of the target volume/lesion 444 relativeto the position and orientation in the planning image. As describedabove, translations may be corrected by translation of thecomputer-controlled treatment table 443, and rotations may be correctedby selection of an appropriate plan from the constrained plan set. Thetranslation of the lesion 444 observed in the cone beam computerizedtomography image relative to the planning image is corrected bytranslation of the patient 441 on the treatment table 443 in the yand/or z directions, and/or by rotation about the x axis. Theorientation of the lesion 444 (i.e., rotations about the y and/or zaxes) are corrected by selecting from the previously calculatedconstrained plan set a modified RTTP that most closely corresponds tothe measured rotation of the lesion 444. Meanwhile, radiographicmonitoring of the patient 441 can be used to check against intrafractionmotion of the patient 441. Furthermore, a cone beam computerizedtomography image acquired immediately prior to, during, or following thetreatment procedure can be obtained in order to provide accuraterepresentation of the location of patient anatomy during treatmentdelivery, which can be stored for off-line review, evaluation, andmodification of subsequent treatment sessions. Following transferal ofthe prescription to the delivery system, the treatment plan is executedaccording to the patient setup and treatment plan determined from thecone beam computerized tomography image.

[0134] The foregoing discussion discloses and describes merely exemplaryembodiments of the present invention. One skilled in the art willreadily recognize from such discussion, and from the accompanyingdrawings and claims, that various changes, modifications and variationscan be made therein without departing from the spirit and scope of theinvention as defined in the following claims. For example, the cone beamcomputerized tomography system can be adapted to perform animal testingidentification, and non-invasive and non-destructive componentstructural testing.

We claim:
 1. A radiation therapy system comprising: a radiation sourcethat moves about a path and directs a beam of radiation towards anobject; a cone-beam computer tomography system comprising: an x-raysource that emits an x-ray beam in a cone-beam form towards said object;an amorphous silicon flat-panel imager receiving x-rays after they passthrough the object, said imager providing an image of said object; and acomputer connected to said radiation source and said cone beamcomputerized tomography system, wherein said computer receives saidimage of said object and based on said image sends a signal to saidradiation source that controls said path of said radiation source. 2.The radiation therapy system of claim 1, wherein said x-ray sourcecomprises a kV x-ray source.
 3. The radiation therapy system of claim 1,wherein said kV x-ray source emits x-rays with energies of approximately100 kV.
 4. The radiation therapy system of claim 1, wherein said x-raysource comprises a linear accelerator.
 5. The radiation therapy systemof claim 1, a stage that moves said object relative to said x-ray sourceand said amorphous silicon flat-panel imager.
 6. The radiation therapysystem of claim 5, wherein said stage rotates about an axis of rotationsaid object relative to said x-ray source and said amorphous siliconflat-panel imager.
 7. The radiation therapy system of claim 2, a stagethat moves said object relative to said x-ray source and said amorphoussilicon flat-panel imager.
 8. The radiation therapy system of claim 7,wherein said stage rotates about an axis of rotation said objectrelative to said x-ray source and said amorphous silicon flat-panelimager.
 9. The radiation therapy system of claim 4, a stage that movessaid object relative to said x-ray source and said amorphous siliconflat-panel imager.
 10. The radiation therapy system of claim 9, whereinsaid stage rotates about an axis of rotation said object relative tosaid x-ray source and said amorphous silicon flat-panel imager.
 11. Theradiation therapy system of claim 1, wherein said x-rays from said x-raysource are emitted along a source plane.
 12. The radiation therapysystem of claim 6, wherein said x-rays from said x-ray source areemitted along a source plane that is perpendicular to said axis ofrotation.
 13. The radiation therapy system of claim 10, furthercomprising an alignment laser that allows visualization of said axis ofrotation and said source plane.
 14. The radiation therapy system ofclaim 1, wherein said amorphous silicon flat-panel imager comprises anarray of individual detector elements.
 15. The radiation therapy systemof claim 14, wherein said array is a two-dimensional array.
 16. Theradiation therapy system of claim 14, wherein each of said individualdetector elements comprises a-Si:H photodiode.
 17. The radiation therapysystem of claim 16, wherein each of said individual detector elementsfurther comprises a transistor coupled to said Si:H photodiode.
 18. Theradiation therapy system of claim 1, wherein said computer receives saidimage from said amorphous silicon flat-panel imager and generates acomputer tomography image of said object based on said received image.19. The radiation therapy system of claim 1, wherein said image is a twodimensional projection image.
 20. The radiation therapy system of claim19, wherein said computer receives said two dimensional projection imagefrom said amorphous silicon flat-panel imager and generates a computertomography image of said object based on said two dimensional projectionimage.
 21. The radiation therapy system of claim 1, further comprising agantry with a first arm and a second arm, wherein said x-ray source isattached to said first arm and said amorphous silicon flat-panel imageris attached to said second arm.
 22. The radiation therapy system ofclaim 21, wherein said gantry rotates about an axis of rotation.
 23. Theradiation therapy system of claim 22, wherein said gantry rotates abouta second axis of rotation.
 24. The radiation therapy system of claim 21,wherein said gantry is attached to a wall of a room.
 25. The radiationtherapy system of claim 23, wherein said gantry is attached to a wall ofa room.
 26. The radiation therapy system of claim 22, wherein saidgantry is attached to a mobile platform that can translationally move ona floor of a room.
 27. The radiation therapy system of claim 23, whereinsaid gantry is attached to a mobile platform that can translationallymove on a floor of a room.
 28. The radiation therapy system of claim 1,wherein said radiation source operates at a power level higher than thatof said x-ray source, wherein said radiation is of an intensity andenergy that is effective for radiation treatment of an area of saidobject.
 29. The radiation therapy system of claim 21, wherein saidradiation source operates at a power level higher than that of saidx-ray source, wherein said radiation is of an intensity and energy thatis effective for radiation treatment of an area of said object.
 30. Theradiation therapy system of claim 1, wherein said x-ray source iscoincident with said radiation source.
 31. The radiation therapy systemof claim 1, wherein said x-ray source is displaced relative to saidradiation source.
 32. The radiation therapy system of claim 1, whereinoperation of said cone beam computerized tomography system with anexternal trigger that controls a biological process of a patient inwhich said object is located.
 33. The radiation therapy system of claim32, wherein said external trigger comprises an active breathing controlmechanism.
 34. The radiation therapy system of claim 32, wherein saidexternal trigger comprises a cardiac gating mechanism.
 35. The radiationtherapy system of claim 1, further comprising an imaging devicepositioned opposite said radiation source and generating an image ofsaid object based on radiation from said radiation source that passesthrough said object.
 36. An imaging system comprising: an x-ray sourcethat emits x-rays towards an object; an imager that receives x-rays fromsaid object based on said emitted x-rays and forms an image of saidobject; an imager support system that attaches said imager to a supportstructure, wherein said imager support system comprises: a first armhaving one end attached to said imager and another end attached to saidsupport structure; and a second arm having one end attached to saidimager and another end attached to said support structure.
 37. Theimaging system of claim 36, wherein said imager support system comprisesa third arm having one end attached to said imager and another endattached to said support structure.
 38. The imaging system of claim 37,wherein said third arm lies in a plane that bisects a line segment thatjoins said one ends of said first and second arms.
 39. The imagingsystem of claim 38, wherein said imager is symmetrically positioned withrespect to said plane.
 40. The imaging system of claim 38, wherein saidimager is asymmetrically positioned with respect to said plane.
 41. Theimaging system of claim 38, further comprising a motorized system thatmoves said imager from a position where said imager is symmetricallypositioned with respect to said plane to a position where said imager isasymmetrically positioned with respect to said plane.
 42. The imagingsystem of claim 37, wherein each of said another ends of said first,second and third arms are attached to a pivot which is attached to saidsupport structure.
 43. The imaging system of claim 36, wherein saidsupport structure comprises a rotating drum of a gantry.
 44. The imagingsystem of claim 43, wherein said x-ray source is attached to saidrotating drum.
 45. The imaging system of claim 44, wherein said x-raysource translates in a direction that is parallel to an axis of rotationof said drum.
 46. The imaging system of claim 43, further comprising aradiation source attached to said rotating drum.
 47. The imaging systemof claim 36, wherein said x-rays emitted from said x-ray source areemitted in a cone beam form.
 48. The imaging system of claim 36, whereinsaid imager comprises an amorphous silicon flat-panel imager.
 49. Animaging system comprising: an x-ray source that emits x-rays towards anobject; an imager that receives x-rays from said object based on saidemitted x-rays and forms an image of said object; an imager supportsystem that attaches said imager to a support structure, wherein saidimager support system comprises: a pivoting arm that has one endpivotably attached to said support structure and another end pivotablyattached to said imager.
 50. The imaging system of claim 49, whereinsaid support structure comprises a rotating drum of a gantry.
 51. Theimaging system of claim 50, wherein said x-ray source is attached tosaid rotating drum.
 52. The imaging system of claim 51, wherein saidx-ray source translates in a direction that is parallel to an axis ofrotation of said drum.
 53. The imaging system of claim 50, furthercomprising a radiation source attached to said rotating drum.
 54. Theimaging system of claim 49, wherein said x-rays emitted from said x-raysource are emitted in a cone beam form.
 55. The imaging system of claim49, wherein said imager comprises an amorphous silicon flat-panelimager.
 56. An imaging system comprising: an x-ray source that emitsx-rays towards an object; an imager that receives x-rays from saidobject based on said emitted x-rays and forms an image of said object;an imager support system that attaches said imager to a supportstructure, wherein said imager support system comprises: a C-armattached to a support structure, wherein said imager is attached to oneend of said C-arm.
 57. The imaging system of claim 56, wherein saidC-arm moves along an arc.
 58. The imaging system of claim 56, whereinsaid support structure comprises a rotating drum of a gantry.
 59. Theimaging system of claim 58, wherein said x-ray source is attached toanother end of said C-arm.
 60. The imaging system of claim 58, furthercomprising a radiation source attached to said rotating drum.
 61. Theimaging system of claim 56, wherein said x-rays emitted from said x-raysource are emitted in a cone beam form.
 62. The imaging system of claim56, wherein said imager comprises an amorphous silicon flat-panelimager.
 63. A method of treating an object with radiation, comprising:move a radiation source about a path; direct a beam of radiation fromsaid radiation source towards an object; emitting an x-ray beam in acone beam form towards an object; detecting x-rays that pass throughsaid object due to said emitting an x-ray beam with an amorphous siliconflat-panel imager; generating an image of said object from said detectedx-rays; and controlling said path of said radiation source based on saidimage.
 64. The method of claim 63, wherein x-rays within said x-ray beamhave an energy of approximately 100 kV.
 65. The method of claim 63,comprising rotating about an axis of rotation said object relative tosaid x-ray source and said amorphous silicon flat-panel imager.
 66. Themethod of claim 63, wherein said amorphous silicon flat-panel imagercomprises an array of individual detector elements.
 67. The method ofclaim 66, wherein said array is a two-dimensional array.
 68. The methodof claim 66, wherein each of said individual detector elements comprisesa-Si:H photodiode.
 69. The method of claim 66, wherein said generatingcomprises forming a computer tomography image of said object based onsaid detected x-rays.
 70. The method of claim 65, further comprisingrotating about a second axis of rotation said object relative to saidx-ray source and said amorphous silicon flat-panel imager.
 71. Themethod of claim 63, further comprising emitting a second set of x-rays,separate from said x-rays emitted from said x-ray source, that have anintensity and energy that is effective for radiation treatment of anarea of said body.
 72. The method of claim 71, wherein said second setof x-rays has an intensity and energy greater than said x-rays emittedfrom said x-ray source.
 73. The method of claim 69, further comprisingcorrecting for offset and gain prior to said generating.
 74. The methodof claim 63, wherein said object comprises an animal.
 75. The method ofclaim 63, wherein said image delineates soft tissue within said animal.76. The method of claim 75, wherein said soft tissue is selected fromthe group consisting of fat, a muscle, a kidney, a stomach, a bowel anda liver.
 77. The method of claim 65, wherein said image is formed afterone rotation of said body relative to said x-ray source and saidamorphous silicon flat-panel imager.
 78. The method of claim 63, whereinsaid x-ray beam is generated by an x-ray source that moves independentlyof said amorphous silicon flat-panel imager, said x-ray source moves ona sinusoidal or sawtooth path constrained to a surface of a cylinderwhile said amorphous silicon panel imager moves in a circular path on asurface of a cylinder.
 79. The method of claim 78, further comprisingadjusting a collimator in real time to adjust a shape of said x-ray beamso it is confined to an active area of said amorphous silicon flat panelimager.
 80. The method of claim 63, wherein said x-ray beam is generatedby an x-ray source that moves dependently of said amorphous siliconflat-panel imager, said x-ray source and said amorphous siliconflat-panel imager each moves on a sinusoidal trajectory on a sphericalsurface.
 81. A method of adding an auxiliary imaging system to anexisting radiation therapy system, said method comprising: providing anexisting radiation therapy system that comprises a radiation source thatis supported on a support structure; and attaching an imager that doesnot directly face said radiation source to said support structure. 82.The method of claim 81, wherein said attaching comprises: attaching saidimager to an imager support system; forming an opening in said supportstructure; inserting a male member through an opening formed in saidimager support system and said opening formed in said support structure;and attaching said inserted male member to said support structure andsaid imager support system.
 83. The method of claim 82, wherein saidattaching said inserted male member comprises tightening a nut onto saidmale member.
 84. The method of claim 81, wherein said support structurecomprises a rotating drum.
 85. The method of claim 82, wherein saidsupport structure comprises a rotating drum.
 86. The method of claim 81,further comprising attaching an x-ray source to said support structure.87. The method of claim 84, further comprising attaching an x-ray sourceto said rotating drum.
 88. The method of claim 85, further comprisingattaching an x-ray source to said rotating drum.
 89. A method ofdelineating a target volume located within a body and shown in acomputerized tomography image, comprising: forming a computerizedtomography image of a target volume in a body; and manually localizingsaid target volume in a slice of said computerized tomography image. 90.The method of claim 89, wherein said manually localizing comprisescontouring of structures in said slice.
 91. The method of claim 89,wherein said manually localizing comprises contouring of structures ofsaid slice in combination with one or more different slices of saidcomputerized tomography image.
 92. A method of delineating a targetvolume located within a body and shown in a computerized tomographyimage, comprising: forming a computerized tomography image of a targetvolume in a body; detecting said target volume identifying translationsand/or rotations of said target volume relative to a position andorientation in a planning image; adjusting the location and/ororientation of said target volume; selecting a radiation therapytreatment plan from a previously formed set of radiation therapytreatment plans as a function of said detected location and/ororientation of said target volume.
 93. A method of delineating a targetvolume located within a body and shown in a computerized tomographyimage, comprising: forming a computerized tomography image of a targetvolume in a body; identifying translations and/or rotations of saidtarget volume relative to a position and orientation in a planningimage; adjusting the location and/or orientation of said target volume;calculating a radiation therapy treatment plan based on said adjustedlocation and/or orientation of said target volume.